Thin strut stent from bioabsorbable polymer with high fatigue and radial strength and method to manufacture thereof

ABSTRACT

This invention discloses method of manufacture of balloon expandable stent made from bioabsorbable polymer with thin struts (strut thickness 130 μm or less, preferably 100-110 μm) with high fatigue and radial strength. The invention further discloses balloon expandable stent made from bioabsorbable polymer with thin struts (strut thickness 130 μm or less, preferably 100-110 μm) with high fatigue and radial strength.

FILED OF THE INVENTION

This invention relates to method of manufacture of balloon expandablestent made from bioabsorbable polymer with thin struts (strut thickness130 μm or less, preferably 100-110 μm) with high fatigue and radialstrength. The invention further relates to balloon expandable stent madefrom bioabsorbable polymer with thin struts (strut thickness 130 μm orless, preferably 100-110 μm) with high fatigue and radial strength.

BACKGROUND OF THE INVENTION

Stents are used to treat atherosclerotic stenosis or other type ofblockages in body lumen like blood vessels or to expand the lumen thathas narrowed due to disease. “Stenosis” is narrowing of the diameter ofa bodily passage or orifice due to formation of plaque or lesion. Thefunction of the stent is to expand the lumen diameter by pressing theplaque to the vessel wall and to maintain patency of the lumen of theblood vessel thereafter at the location of its implantation. The stentmay be coated with therapeutic agent/s and/or biocompatible material/sfor beneficial effects like minimizing the possibility of restenosis,reduction in inflammation etc.

The first step in treatment of stenosis involves locating the regionthat may require treatment such as a suspected lesion in a vessel byangiography of the diseased vessel followed by implanting a suitablestent. The stent may be balloon expandable type or self-expanding type.The stents are mounted on the delivery catheter which helps indelivering the stent to the target site of the disease.

The balloon expandable stent is mounted on a balloon catheter bycrimping process such that it holds tightly over the balloon and attainsa considerably lower diameter (profile). The catheter is percutaneouslyinserted into the body lumen and is directed to the site of the disease(blockage or narrowed lumen). At the site of the disease, the balloon isinflated by application of hydraulic pressure to expand the stentradially to desired diameter. Radial expansion of the stent presses theplaque to the wall of the vessel by which the restriction to the flow ofblood in the vessel is removed. The balloon is then deflated by removingthe hydraulic pressure and withdrawn from the body of the patient.

On expansion, the stent material attains plastic deformation and hencethe stent does not′ recoil back to its original shape and remains inexpanded state keeping the lumen patent. Self-expanding stents aretypically made of metal with shape memory and they expand without helpof any other device such as balloon. The stent is mounted on thedelivery catheter and expansion of the stent is restricted by a sheath.The catheter is percutaneously inserted into the body lumen and guidedto the target site where the lesion or plaque is located. The sheath isthen retracted to allow the stent to expand. Like balloon expandablestent, this stent also makes the lumen patent by pressing the plaque.

The structure of stents is cylindrical with scaffold made up of apattern or network of interconnecting structural elements i.e. “struts”.The scaffolding of the stent may be formed from wires, tubes, or sheetsof material rolled into a cylindrical shape. In addition, the surface ofthe stent may be coated with formulation of therapeutic agent/s and/orbiocompatible materials with suitable carriers and additives.

It is important that the stent must be able to withstand structuralloads viz. radial compressive forces imposed by the wall of the bodylumen on the stent. Radially directed force from the wall of the lumenmay tend to cause the stent to recoil inward. The radial strength of thestent must be adequate to resist radial compressive forces. These forcesare cyclic in nature due to pulsating blood flow. Hence the stent shouldhave adequate fatigue strength to withstand cyclic loading imposed on itby the lumen. In addition, the stent must possess sufficient flexibilityto allow for crimping, maneuvering through the vascular pathway, andexpansion process. The scaffold structure should also be dense enough toprevent prolapse of the plaque but open enough to allow easy side branchaccess for another catheter with or without stent. The stent shouldexhibit required radio opacity for ease of implantation.

Stents have been used effectively for quite a long time and the safetyand efficacy of stenting procedure are well established. Implantation ofthe stent causes some injury to the vessel. The healing process startsand finally the endothelial cells are formed at the implantation site.Once the healing process is completed the endothelial cells providesufficient support to the wall of the lumen and the stent is no longerrequired. Thus, the presence of the stent in the lumen is required onlyfor a limited period of time till the healing process is completed.

Coronary stents are generally made from biocompatible materials such asmetals which are bio-stable. Metal has high mechanical strength thatprovide adequate radial and fatigue strength to the stent that preventearly and later recoil. However, the metallic stent remains at theimplant site indefinitely. Leaving the stent at the implanted sitepermanently causes compliance difference in the stented segment and thehealthy vessel segment. In addition, there is possibility of permanentinteraction between the stent and the surrounding tissue resulting in arisk of endothelial dysfunction causing delayed healing and latethrombosis.

Drug-eluting stents are a breakthrough in the development of stents withtheir ability to significantly reduce restenosis rates and the need forrepeat revascularization. However, they are still associated withsub-acute and late thrombosis that necessitates prolonged antiplatelettherapy for at least 12 months.

Metallic stents have been used effectively for quite a long time andtheir safety and efficacy are well established. The main issues of astent are restenosis and in-stent thrombosis. One of the importantcauses of these adverse effects is injury of the artery caused byimplantation of the stent. The injury leads to restenosis and delayedendothelialization. These adverse effects can be reduced if injury tothe artery is reduced.

It is well established that the thickness of the struts of a stent playsan important role in injuring of the artery. Thinner struts cause lessinjury compared to thicker struts. Thus, the injury of the artery can bereduced by making the struts as thin as practically possible. Whiledeciding the thickness of the struts, care should be taken so thatimportant mechanical properties of the stent like radial strength andfatigue resistance are adequate to withstand forces imposed by the bodylumen like artery.

The injury to the artery wall can thus be minimized by reducing thethickness of the struts of the stent scaffold structure. It is wellestablished that the stent with less strut thickness causes less injurycompared to the stent with thicker struts. This subject is discussed indetail by Kastrati A, Schömig A, Dirschinger J, et al. in their paper“Strut Thickness Effect on Restenosis Outcome (ISAR STEREO Trial)”published in Circulation 2001; 103:2816-2821. The incidence ofangiographic restenosis was 15.0% in the group of patients treated withstents of thin struts against restenosis of 25.8% in the group treatedwith stents with thicker struts. Clinical restenosis was alsosignificantly reduced, with a reintervention rate of 8.6% amongthin-strut patients and 13.8% among thick-strut patients.

These findings were reconfirmed by Kastrati A, et al in their paper“Strut Thickness Effect on Restenosis Outcome (ISAR STEREO-2 Trial)”published in J. Am. Coll. Cardiol, 2003; 41:1283-8. The incidence ofangiographic restenosis was 17.9% in the group of patients treated withstents of thin struts against restenosis of 31.4% in the group treatedwith stents with thicker struts. Target Vessel Revascularization (TVR)due to restenosis was required in 12.3% of the patients in thin strutgroup against 21.9% required in patients of the thick strut group.

In conclusion from above it was established that the use of thinnerstrut device is associated with a significant reduction in angiographicand clinical restenosis after coronary stenting.

The stents can be made from polymeric materials which arebio-absorbable/biodegradable. A biodegradable stent can be configured todegrade and disappear from the implant site when it is no longer neededleaving behind only the healed natural vessel. This will allowrestoration of vasoreactivity with the potential of vessel remodeling.These stents are believed to improve the healing process whereby thechances of late stent thrombosis are reduced considerably. Prolongedantiplatelet therapy then may not be necessary. The biodegradable stentsmay be made from biocompatible polymers such as Poly-L-lactic acid(PLLA), polyglycolic acid (PGA), poly (D,L-lactide/glycolide) copolymer(PDLA), and polycaprolactone (PCL). Poly-L-lactic acid (PLLA) is usuallyrecommended polymer among others.

The only disadvantage of polymeric materials is their lower mechanicalstrength compared to the metals. Strength to weight ratio of a polymericmaterial is smaller than that of a metal. This makes it necessary toincrease thickness of the polymeric stents compared to metallic stentsto get adequate radial and fatigue strengths. The increase in thicknessresults into higher profile and higher degree of injury to the bloodvessel. Higher thickness reduces the flexibility of the stent resultinginto poor trackability through tortuous arteries. Polymeric materialshave poor radio opacity. Polymeric materials are also brittle underconditions within human body.

It is hence necessary to select right polymer and modify its mechanicalproperties to make it suitable for stent application. Making the stentwith low strut thickness poses additional challenge. Selection of apolymeric material, design of stent scaffold structure and process formaking a stent require careful attention to several aspects. The stentshould have adequate mechanical strength to prevent recoil. The rate ofdegradation of the polymer should be such that the mechanical strengthof the stent is retained to provide support to the vessel and preventprolapse of the plaque into the vessel till the healing process iscomplete. The stent should eventually disappear by degradation. Thestent should have enough flexibility for ease of crimping on the balloonof the catheter and for good trackability through the tortuous passagesthrough arteries. The polymeric material and its degradation productsshould be biocompatible. The rate of degradation will influence therelease profile of the therapeutic agents coated on the stent. Polymerssuch as Poly-L-lactic acid (PLLA), polyglycolic acid (PGA), poly(D,L-lactide/glycolide) copolymer (PDLA), and polycaprolactone (PCL) andtheir degradation products are known to be non-toxic and biocompatible.

There is a continuing need for manufacturing and fabricating methods forpolymeric stents with such scaffold design that offer adequate radialstrength, fracture toughness, low recoil and sufficient shape stabilitywith low strut thickness. A stent with low strut thickness will resultin low injury to the arterial wall. In addition, thin stent will givelower profile in crimped condition compared to a stent with higher strutthickness. Stents with thinner struts impart more flexibility to thestent.

There is ample literature available on biodegradable stents and theprocess for manufacturing the same.

U.S. Pat. No. 7,971,333 describes method of forming a stent frompolymeric materials by modifying the mechanical properties of polymertube to get desirable mechanical properties. The polymer can be modifiedto increase the strength, modulus and/or toughness of the polymer tubeto make them comparable to metal. Mechanical properties of a polymer canbe modified by applying stress to the polymer preferably above its glasstransition temperature (Tg) followed by heat setting. This inducesmolecular orientation of polymer chains in radial and axial directions.The stress is applied to the polymer tube by expanding it radially byblow molding and by stretching the tube axially by applying axial loadthat result into biaxial orientation of polymer molecules. The tube isheated to desired temperature by heating the mold. Radial deformation ofthe tube is achieved by pressurizing the tube in the mold with inert gasunder pressure. The degree of radial deformation is defined as ratio ofoutside diameter of tube after expansion and original internal diameterof tube. This ratio may vary between 1 and 20 or narrowly between 2 and6. Degree′ of axial deformation is defined as ratio of lengths of tubeafter and before deformation. Temperature and degree of deformationaffect crystallinity which in turn is dependent on crystallinity of thetube before deformation. The patent describes laser cutting of thedeformed tube to get the scaffold structure of the stent.

U.S. Pat. No. 8,501,079 discloses method for fabricating a stent fromPLLA tube; radially and axially expanding the tube inside a mold whilethe tube is heated to a processing temperature; wherein the processingtemperature is 84° C. The radial and axial expansion percentages are400% and 20% respectively to produce an expanded tube having anincreased mechanical strength, fracture toughness and homogeneity in amechanical properties over the wall thickness of the expanded tube; andforming the stent from the expanded tube. The radial expansion of thetube is achieved at a pressure of 110-140 psi.

US 2013/0187313 discloses a method for fabricating stent comprisingproviding a PLLA tube disposed within a cylindrical mold; heating themold and the tube to a tube deformation temperature (80° C. to 115° C.)with a′ heat source translating along the cylindrical axis of the moldand the tube, wherein the heat source translation rate is between0.2-1.2 mm/sec; increasing the pressure inside the tube; allowing theincreased pressure in the tube (110-140 psi) to radially expand the tubeagainst the inner surface of the mold, wherein the radial expansionpropagates along the cylindrical axis of the mold and tube as the heatsource translates along the cylindrical axis, applying a tensile forceto the tube along the cylindrical axis during the radial expansion toaxially elongate the tube during the radial expansion, wherein thepercent radial expansion is 300-500% and the percent axial elongation is100-200%; and forming a stent pattern in the axially expanded andradially deformed tube.

EP1973502 reports a stent comprising a deformed spherical radiopaquemarker disposed in a depot in a portion of the stent, the marker beingcoupled to the portion at least partially by an interference fit betweenan expanded portion of the marker and an internal surface of the portionof the stent within the depot, wherein the marker comprises sufficientradio opacity for ease of imaging by normal imaging techniques. Gapsbetween the deformed marker and the internal surface are filled with apolymeric coating material.

US 2011/0066222 describes method of forming a stent from PLLA tubularpolymer that is deformed in a blow mold. Desired polymer morphologyresulting in improved stent performance is obtained with axial expansionratio from about 10%-200%, preferably 20% to 70%, a radial expansionratio from about 100%-600%, preferably 400% to 500%, axial deformationpropagation at or about 0.3 mm/min, selected expansion pressure of about50 to 200 psi, preferably 130 psi and expansion temperature about 100°F. to 300° F. preferably less than 200° F. Heating is done by a movingheat source outside the mold. The heat source is moved at the rate of0.1-0.7 mm per minute. The stent may be made of PLGA, PLLA-co-PDLA,PLLD/PDLA stereo complex, and PLLA based polyester block co-polymercontaining a rigid segment of PLLA or PLGA and a soft segment of PLC orPTMC.

None of the prior art mentions the design and manufacturing method formaking polymer stent with low strut thickness (less than 130 μm,preferably 100-110 μm thickness).

The polymer stents have potential shortcomings compared to metal stentsof the same dimensions viz. lower radial strength and lower rigidity ofthe polymer stents compared to metallic stents. Lower radial strengthpotentially contributes to relatively high recoil of polymer stentsafter implantation into an anatomical lumen. Another potential problemwith polymer stents is that struts can crack or fracture duringcrimping, delivery and deployment, especially for brittle polymers. Dueto these shortcomings, the strut thickness of the polymer stents isalways kept higher compared to the metallic stents with same radial andfatigue strength.

In conclusion from “ISAR STEREO Trial” and “ISAR STEREO-2 Trial” asdescribed above, it was established that the use of thinner strut deviceis associated with a significant reduction in angiographic and clinicalrestenosis after coronary stenting.

Thus, there is a continuing need for identifying right polymer andmanufacturing and fabricating methods for polymeric stents of rightscaffold design that impart sufficient radial strength, fracturetoughness, low recoil and sufficient shape stability at low strutthickness. Additional advantage of a stent with low strut thickness isits lower profile after it is crimped on the balloon of the deliverycatheter and more flexibility.

Making a stent with desired low strut thickness starts with choosing aright polymeric material. The polymeric material then undergoes a numberof process steps like drawing the tube from this material, modifying themechanical properties of the tube, making the stent from this tube withright scaffold design, crimping the stent on the balloon of the deliverycatheter and sterilization of the assembly.

The tube of the chosen polymer may be formed by extrusion or moldingprocess under controlled conditions to achieve desired properties of thetube. Processing conditions that affect tube properties mainly includedraw down ratio during extrusion, temperature at which the tube isextruded (relative to glass transition temperature and melting point ofthe polymer) and tube diameter.

Mechanical properties of a polymer can be modified by applying a stress.The stress alters the molecular structure and/or morphology of thepolymer. The degree and rate of changes in mechanical properties dependon the temperature at which the stress is applied and degree ofdeformation the polymer (the tube in this case) undergoes due toapplication of the stress. The stress can be applied to the polymer tubein radial and axial directions to modify the crystalline morphology andpolymer chain orientation in controlled manner to achieve a desiredcombination of strength and fracture toughness along axial and radialdirections. Combined with right scaffold design, the strut thickness canbe reduced while maintaining high fatigue and radial strength andkeeping recoil under control. At the same time, it is necessary toachieve desired degradation rate of the polymer such that the stentretains adequate mechanical strength till the healing process of thelumen is completed and the stent eventually disappears from the implantsite. The processing of the tube in this way changes the crystallinityof the polymer which in turn influences the degradation rate of thepolymer. Amorphous polymer degrades faster than crystalline polymer butit is mechanically weaker than the crystalline polymer. Hence, a balanceis required to be achieved in processing of the tube such that the stenthas right combination of mechanical strength and degradation rate.

In the light of the above there is a need in the art to develop abiodegradable polymer stent with thin struts from a bioabsorbablepolymer with adequate fatigue and radial strengths and a method tomanufacture thereof. The process begins with choosing right grade ofpolymer and setting extrusion process to get desired properties ofextruded tube. The grade of the polymer is characterized broadly by itsmolecular weight, glass transition temperature (T_(g)), crystallinity(X_(c)), molecular structure, and stereo isomerism. Further processingof the tube into stent and the scaffold design of the stent structureshould be such as to achieve desired mechanical properties of thefinished stent. The processing of the tube includes application ofstress to the tube, laser cutting, cleaning of the cut stent, radioopaque marker deposition, heat treatment, drug coating, crimping andsterilization.

The mechanical properties are largely dependent on polymercharacteristics like average molecular weight and molecular weightdistribution. These characteristics undergo change at each processingstage. Hence it is necessary to check these characteristics at eachprocess stage and devise a process that results in high mechanicalstrength of the finished stent.

Sterilization of the stent is done by e-beam radiation and this steprequires special attention. The e-beam radiation causes degradation ofpolymer and thus has a significant effect on the average molecularweight of the bioabsorbable polymer and hence its mechanical properties.The current inventors have studied the effect of e-beam radiation onpolymer over a wide range of e-beam doses and found that reduction inthe e-beam dose improves the mechanical strength of the polymer. Normaldose of e-beam for effective sterilization is more than 20 kGy. The dosecan be reduced to some extent by adding stabilizer/s in the polymermatrix. This stabilizer should be biocompatible and should not createany adverse clinical effect.

Therefore, one of the objectives of the instant invention is to geteffective sterilization with e-beam dose considerably lower than 20 kGywithout use of any additive.

Accordingly, the objective of the invention is to provide abiodegradable/bioabsorbable stent made from a bioabsorbable polymer withthin struts (thickness 130 μm and less, preferably with thickness of100-110 μm) which has adequate fatigue strength, radial strengths andlow recoil and a method for manufacture thereof, for which protection issought.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a view depicting the mold system where the polymer tube isprocessed.

FIG. 2 depicts details of the end plug to hold the tube.

FIG. 3 depicts polymer tube inside the mold under process of radialdeformation.

FIGS. 4, 4A, 4B, 4C, 5, 6 and 7 depict the scaffold structures of thestent.

FIG. 8 depicts the shape of the radio-opaque marker.

SUMMARY OF THE INVENTION

The terms “bioabsorbable” and “biodegradable” are used interchangeablythroughout the specification and the same may be appreciated as such bythe person skilled in the art.

In accordance with the above objectives, the present invention disclosesa process for preparation of biodegradable polymer stent made of PLLA(poly-L-Lactide) with strut thickness of less than 130 μm whichcomprises:

-   -   (a) Deforming the extruded PLLA tube having weight average        molecular weight M_(w) in the range of 590000 to 620000, number        average molecular weight M_(n) in the range of 350000 to 370000        and crystallinity in the range of 7% to 12%, axially at 70°        C.-80° C. by applying axial force till desired stretch is        achieved and radially expanding the tube at a temperature of        70° C. to 80° C. by pressurizing the tube with inert gas in        three stages viz., 250-280 psi in stage-1, 375-410 psi in        stage-2 and 500-530 psi in stage-3;    -   (b) Heating the tube after radial deformation under the same        pressure conditions between 100° C. and 110° C. and maintaining        for up to 2 min and then cooling to 20° C. in 20-30 sec to get        finished deformed tube;    -   (c) Cutting specific pattern of scaffold structure on the        deformed tube by laser machining;    -   (d) Annealing the laser cut stent before or after depositing the        radio opaque markers;    -   (e) Cleaning the annealed stent with radio opaque markers using        solvent to remove irregularities and to achieve smooth surface;    -   (f) Coating the cleaned stent with a formulation of        antiproliferative drug and a career polymer by spray coating        method;    -   (g) Crimping the coated stent on the balloon of pre-sterilized        delivery catheter in clean environment;    -   (h) Sterilizing the crimped stent and catheter system by e-beam        method with e-beam dose less than 20 kGy without compromising        effective sterilization.

The axial deformation of the tube according to the invention is carriedout at temperature between 70° C. and 80° C., preferably between 74° C.and 76° C. at elongation ratio between 1.4 and 1.7 and maintaining thetemperature and pressure conditions for 15-20 sec.

The radial deformation of the tube in accordance with the invention iscarried out at temperature between 70° C. and 80° C., preferably between74° C. and 76° C. at radial expansion ratio between 3 and 5 bypressurizing the tube with nitrogen in three stages as mentioned in (a)above and maintaining the temperature and pressure conditions for 15-20sec after each stage of pressurization. According to the invention, thetube after radial deformation is heated between 100° C. and 110° C. andmaintained for 30 sec to 2 min and then cooled to 20° C. in 20-30 sec toget finished deformed tube with PDI lower than the extruded tube.

In the process of the present invention, the laser cutting operation ofthe finished deformed tube is done at wave length between 1300 and 1600nm.

In a preferred aspect, the scaffold structure of the stent comprises apattern with rows of struts with sinusoidal shape wherein the peaks ofone row is aligned with valleys of the other row and wherein, the valleyof one row is connected by straight crosslinking struts with peak ofsubsequent row at every third such position in the central portion andat every peak and valley in the end portions.

According to the invention, the annealing is done at temperature between100° C. and 110° C. for a period varying from 3 hours to 4 hours undervacuum of up to 650 to 700 mm Hg followed by cooling the stent toambient temperature in 25-30 sec.

Further, cleaning of the annealed stent is done by rotation in perchloroethylene as such or diluted in a suitable solvent or in a mixture of isopropyl alcohol and chloroform.

In the process of the present invention, six radio opaque markers madefrom platinum with triaxial ellipse shape are fixed on the cross linkingstruts of the end portions of the stent, wherein, three markers arefixed at each end of the stent wherein the said markers are equi-spacedat 120° to each other such that they give clear idea of the stentposition as well as patency of the stent at the ends in two standardorthogonal views without help of OCT or IVUS.

In another preferred aspect, the stent is coated with formulation ofSirolimus with biodegradable polymer viz. PDLLA in 50:50 proportion w/wand Sirolimus dose of 1.25 μg per mm² area of the stent. The coating isdone using spray coating.

In yet another aspect, the invention provides crimping of the coated ornon-coated stent. According to this method, the coated or non-coatedstent is crimped on the pre-sterilized balloon of the delivery catheterunder clean atmosphere at 25° C. to 40° C. in 6-8 stages and dwell timeof 200-310 sec.

In yet another aspect, the stent is sterilized by;

-   -   (a) subjecting the components other than the stent to ETO        sterilization process;    -   (b) crimping the un-sterilized stent on balloon of the        sterilized catheter in clean environment; and    -   (c) subjecting the entire stent system to e-beam sterilization        at dose between 6 and 12 kGy, preferably between 6 and 10 kGy at        temperature between 15° C. and 25° C.        During the sterilization process, the bio burden of less than 3        cfu and Sterility Assurance Level (SAL) of “six log reduction”        (10-6) are achieved with no significant effect on the optical        rotation or crystallinity.

By using the process of the present invention, a stent with strutthickness of 100-110 μm can be achieved with radial strength of 20-25Ndepending on stent size and adequate fatigue strength.

Accordingly, the invention further encompasses balloon expandable stentmade from bioabsorbable polymer with thin struts (strut thickness 130 μmor less, preferably 100-110 μm) with high fatigue and radial strength.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

The present invention discloses a biodegradable stent with thin strutswith adequate fatigue and radial strength as well as low recoil, madefrom tube of right grade of PLLA polymer and a method of manufacturethereof. The various embodiments of the present invention describe thepolymer properties and manufacturing of the stent. Present invention canbe applied to balloon expandable stents, stent grafts or stents forother vascular applications.

The mechanical properties of a polymer depend largely on characteristicslike average molecular weight and molecular weight distribution. Thepolymer has different size and types of monomer chains. Molecular weightof a polymer can be described by weight average molecular weight My, andnumber average molecular weight M_(n). M_(w) represents averagemolecular mass of various molecular chains of the polymer which includeseven those having same types of individual macro molecules of differentchain lengths. M_(n) represents the average of different sizes ofvarious polymer chains and it is arithmetic mean or average of themolecular masses of the individual macromolecules. The M_(w) and M_(n)can be determined by gel permeable chromatography (GPC). Anotherimportant parameter for a polymer is the poly dispersity index (PDI)which is the ratio of M_(w) to M_(n)(M_(w)/M_(n)). This parameter givesan indication of how narrow the molecular distribution is. A parameterclosely related to M_(w) and M_(n) is the intrinsic viscosity which was,measured using Brookfield viscometer model LVDV E230.

In addition, other parameters described further are important for thepolymer. Glass transition temperature T_(g), and melting temperatureT_(m) are important thermal properties. Processing of a polymer atelevated temperatures results in the change in morphology of a polymerand influences its crystallinity X_(c). X_(c) defines the degree ofcrystallinity of a polymer in percentile value. A totally amorphouspolymer has X_(c) value of 0% and a fully crystalline polymer has X_(c)value of 100%. Polymers with higher microcrystalline regions (higherX_(c)) are generally tougher and more impact-resistant than polymerswith lower microcrystalline regions (lower X_(c)).

Since PLLA is the polymeric form of an optically active monomer,specific rotation of PLLA is also an important characteristic. Polymersobtained from optically active monomers are semi crystalline whileoptically inactive monomers give amorphous polymers. Crystalline polymerhas higher mechanical and thermomechanical properties as describedabove. The difference in mechanical properties are related to the stereoregularity of the polymer chains which are characterized by presence ofonly S(-) chiral centers. For example, the propensity for the lactidemonomer to undergo racemization to form meso-lactide can impact opticalpurity and thus material properties of the polymer at highertemperatures.

All these characteristics influence the mechanical properties of thepolymer used for the process of making the stent.

In the present invention, for making a bioabsorbable stent, a number ofbioabsorbable polymers like PLLA, PLGA, PDLA of various molecularweights were studied. It was observed during the study that M_(w) andM_(n) and also other characteristics change at each step of the stentmanufacture i.e. starting with extrusion of the tube till sterilization.Observation of these properties under various processing conditions gavevaluable insights which helped in making the finished stent withadequate mechanical strength in spite of low strut thickness. Specificscaffold design also played an important role in achieving low strutthickness.

The polymer tube is formed using extrusion or molding process. Duringextrusion or molding process, PLLA can undergo thermal degradation,leading to formation of lactide monomers and other by-products whichresult in reduction of molecular weight of the polymer. Dependence ofmechanical properties of PLLA with its molecular weight is wellinvestigated; the strength increases with increase in molecular weight.Degradation of the polymer causes reduction in its average molecularweight. Hence, excessively high temperatures should be avoided duringextrusion to avoid degradation of the polymer. The extrusion of a tubeis done at about melting point of the polymer and significant decreasein average molecular weight M_(w) as well as M_(n) was observed in allthe types of bioabsorbable polymers. It was observed that under similarconditions, % reduction in M_(w) was dependent on the M_(w) beforeextrusion process. This means that % reduction in M_(w) was higher forpolymer with higher M_(w) compared to the polymer with lower Mw. Forexample, M_(w) of PLLA before and after extrusion were found to be765230 and 590780 respectively which amounted to 22.8% reduction inM_(w). Another PLLA with Mw 622480 under similar conditions showed Mw496490 after extrusion i.e. Mw reduced by 20.24%. Similar results havealso been observed in case of PLGA and PDLA tubes.

In both the above cases, the % reduction in average molecular mass M_(n)was higher than M_(w). In the former case, M_(n) reduced from 563340 to355280 i.e. 36.94% and in latter case, M_(n) reduced from 459630 to307720 i.e. 33.05%. Thus, the extrusion conditions can influence theproperties of the polymer in a significant manner. It is also importantto note that mechanical properties of the extruded tube are alsodependent on the stresses applied during extrusion process and the otherprocess parameters.

Mechanical properties of polymer depend on molecular orientation of thepolymer chains. Molecular orientation of polymer chains is altered whenstress is applied to the polymer. The reorientation of the molecularchains occurs in the direction of the applied stress. The extent towhich the orientation of polymer chains gets altered depends on thetemperature at which the stress is applied and the magnitude of thestress. Generally, for altering the molecular orientation, thistemperature should be above glass transition temperature (T_(g)) of thepolymer and lower than its melting point. The stress can be applied inaxial and radial directions to orient the polymer molecules axially andcircumferentially.

The application of the stress also influences crystallinity of thepolymer. As mentioned above, mechanical properties of the polymer alsodepend on its crystallinity. Crystallinity increases the mechanicalstrength of the polymer. Crystallinity also influences the degradationrate of the polymer. Increase in crystallinity decreases the degradationrate. The method of manufacture and scaffold structures described beloware with respect to PLLA as polymer. However, the aspects of thisinvention can be applied to other polymers and one skilled in the artcan adopt the aspects of this invention by optimizing them for differentbiodegradable and biocompatible polymeric materials.

Accordingly, in a preferred embodiment, the invention provides adetailed method for manufacturing a polymer (PLLA) stent as describedbelow:

-   -   (a) The polymer tube is made by extrusion or injection molding.        The process conditions and equipment should produce the tube        with desired internal diameter d_(i) and external diameter        d_(o). In the instant invention, the polymer tube was made by        extrusion process and M_(w) of the polymer tube ranged from        590000 to 620000, M_(n) ranged from 350000 to 370000, and        crystallinity ranged from 7% to 12%.    -   (b) The extruded polymer tube T with internal diameter d_(i) and        external diameter d_(o) is placed centrally in the mold M        (depicted in FIG. 1). Mold M is cylindrical in shape and is made        from metal with good thermal conductivity. The mold M has inside        diameter DM as shown in FIG. 1. The polymer tube T is aligned        centrally in the mold by use of end plugs P provided at both        ends of the mold M. Details of the end plug P are depicted in        FIG. 2. The portion of the polymer tube T inside the mold M is        designated as T′.    -   (c) Two ends (proximal end E1 and distal end E2 as depicted in        FIG. 1) of the polymer tube T are held firmly outside the mold        in fixtures F and C. The distal end of the tube T is fixed to        fixture C such that this end of the tube T gets crimped and        sealed. Both the fixtures (F and C) can be moved by applying        axial force F_(A) equally on both fixtures which gets        transmitted axially to the tube T.    -   (d) The proximal end E1 of the tube T is connected to a source        of inert gas G (as depicted in FIG. 1) which can supply the        inert gas under pressure inside the tube T.    -   (e) The mold M is covered by jacket J (shown in FIG. 1) which        has heating and cooling arrangement. The heating is done        electrically and cooling by suitable cooling medium. The jacket        J does the heating and cooling of the mold M which in turn heats        and cools the polymer tube T. The heating and cooling system is        capable of heating or cooling the mold M and the tube T′        uniformly. High thermal conductivity of the mold helps in        uniform heat transfer for the jacket J, the mold M and the tube        T′ inside the mold. Temperature indicators are provided at        critical locations to exhibit the temperature in the mold M.    -   (f) The electrical heating is turned on in the jacket J and the        tube T′ is heated to temperature between 70° C. and 80° C.,        preferably between 74° C. and 76° C.    -   (g) The tube T is then axially deformed (elongated) by applying        tensile force F_(A) to both the fixtures (F and C). The        temperature of the portion of the tube which is inside the mold        M (T′) is higher than rest of the tube (T). Hence, this portion        will start deforming axially i.e. its length will increase. The        axial force is applied till desired elongation of tube T′ is        achieved. The ratio of final length of the tube to its initial        length is termed as “axial expansion ratio”. This ratio is kept        between 1.4 and 1.7.    -   (h) The conditions in the mold (tensile force and the        temperature) at this stage are held for 15-20 sec to set the        tube condition. Then the tensile force is removed.    -   (i) Inert gas like nitrogen under pressure from the inert gas        source G is then introduced in the tube through its proximal end        E1 while maintaining the temperature. As the distal end E2 of        the tube T is crimped and sealed at C, the pressure in the tube        T/T′ will increase. The temperature of the portion of the tube        which is inside the mold M (T′) is higher than rest of the tube        (T). Hence, this portion will start deforming radially i.e. its        diameter will increase. The inert gas pressure is applied in two        or more stages, preferably in three stages. The stage wise        application of inert gas pressure improves the mechanical        properties further, reduces recoil considerably and eliminates        possibility of formation of cracks and waviness on the surface        of deformed tube T′. The external diameter of the deformed tube        T′ will increase till the external surface of the tube contacts        internal surface of the mold M as depicted in FIG. 3. The inside        diameter (DM) of the mold M restricts the extent of deformation        of tube T′. The internal and external diameters of the tube T′        will increase to D_(i) and D_(o) respectively as shown in        FIG. 3. The ratio of internal diameter of expanded tube D_(i) to        original tube diameter d_(i) is termed “radial deformation        ratio”. D_(o) is dependent on the internal diameter of the mold        D_(M). Hence, the internal diameter of mold M (D_(M)) is kept        such that the desired radial deformation ratio is achieved. The        radial deformation ratio is kept between 3 and 5.

The inert gas pressure and the temperature in the Mold M are maintainedfor 15-20 sec after each stage to set the condition.

While maintaining the inert gas pressure in the tube, the temperature ofthe tube is raised and maintained for a specific period of time toachieve consistent properties of the tube. Cooling medium is thenintroduced in jacket J to cool the mold M and the deformed tube T′ whichis then removed from the mold M.

-   -   (j) The deformed tube is then cut on the laser machine to form        the stent scaffold structure on the tube.    -   (k) Radio opaque markers are then fixed on the stent.    -   (l) The laser cut stent with radio opaque markers is then        annealed under vacuum to get consistent polymer morphology        throughout the scaffold. This step helps to remove residual        monomers and reduced lot to lot variation of the properties. The        aim of this step is not to increase crystallinity. There is very        little change in crystallinity of the polymer during this step.    -   (m) The stent is then cleaned with suitable solvent to remove        any surface irregularities and achieve smooth surface, followed        by removal of solvent under vacuum.    -   (n) The stent is then coated with the formulation of therapeutic        agent like antiproliferative drug.    -   (o) The stent is then crimped on the delivery device viz.        catheter between 25 and 40° C.    -   (p) The crimped stent along with the delivery system is then        sterilized using e-beam.

Each manufacturing step described above affects polymer properties likemolecular weight, crystallinity, molecular orientation etc. This in turnchanges the mechanical properties of the polymer. The mechanicalproperties of the finished stent should be adequate to demonstratesufficient radial and fatigue strengths as well as low recoil. The stentshould have desired degradation rate so that the stent provides adequatesupport to the blood vessel till it is healed and eventually disappearfrom the implant site.

Due to axial and radial deformations, the length and the diameter of thetube undergo change. Hence, the thickness of the tube will change.Thickness of the deformed tube T′ will be lower than that of theoriginal tube T. To get the stent of desired thickness, the thickness ofthe deformed tube should be controlled by choosing the original tube Twith specific internal and outer diameters d_(i) and d_(o) anddeformation ratios to get desired dimensions of the processed (deformed)tube (D_(i) and D_(o)). These dimensions determine the strut thicknessof the finished stent scaffold.

The tubular mold used for deforming the tube is made of metal like highgrade of beryllium copper which has very good thermal conductivity.

A simple end plug P is used at both ends of the mold to keep the tube atthe center of the mold. Details of plug P are depicted in FIG. 2. Thisleads to uniform heating of the tube which ensures uniform axial andradial expansion of the tube in the mold. This in-turn results inuniform thickness of the deformed tube. For this process, use ofcomplicated arrangement like heat source which translates across themold surface at controlled rate is not required.

In a specific embodiment, inert gas pressure is applied to the tube inthree stages i.e 250-280 psi in stage 1, followed by 375-410 psi instage 2 and finally 500-530 psi in stage 3. Temperature is maintainedbetween 70° C. and 80° C., preferably between 74° C. and 76° C. Stagewise application of such relatively higher pressures ensures deformationof the tube T′ with tight tolerance and elimination of cracks and wavysurface. After each stage, the conditions (pressure and temperature) aremaintained for a period of time varying between 15 and 20 seconds to setthe tube under each of these conditions.

While maintaining the pressure of the last stage i.e. 500-530 psi, thetube is heated to temperature between 90° C. and 120° C., preferablybetween 100° C. and 110° C. and maintained for 30 sec to 2 minutes. Thetube is then cooled to 20° C. in 20-30 sec. The pressure is thenreleased and the tube is then removed from the mold. The tube at thisstage achieves consistent properties. The crystallinity at this stage isless than 45%.

The three stage pressurization of the tube offers advantage over asingle stage. The overall % reduction in M_(w) and M_(n) after annealingwas lower in case of three stage pressurization compared to a singlestage pressurization. The Poly dispersity index PDI was lower in case ofthree stage process compared to single stage which indicated narrowmolecular weight distribution. Three stage pressurization resulted inthe deformed tube with tight tolerance and free of cracks and wavysurface.

In yet another embodiment, laser cutting of the deformed tube is doneusing femto seconds equipment and laser beam of 1300 to 1600 nmwavelength to cut polymer scaffold on the processed tube. The lasercutting process creates the stent scaffold pattern which may consist ofstruts which are structural elements formed on the tube by laser cuttingprocess. The diameter of the tube after radial expansion may be betweeninitial diameter of the tube (d_(i)/d_(o)) and expanded diameter of thefinished stent. The scaffold structure of a stent has repetitiveradially expandable rows of geometrical shapes across its circumferencewhich may be termed as cylindrical elements forming rings. The shape ofan element and the way such elements are interconnected with each othercan be manipulated to achieve different structural properties i.e.mechanical strength which imparts resistance to radial forces applied tothe stent structure by the vascular lumen walls. There exists largedesign flexibility in creating different shapes. This flexibility shouldbe used keeping in mind other desirable properties of the stent. Thestent structure is formed by placement of these elements in a specificpattern to form specific shapes and interconnected array of struts. Theelements in the pattern should be close enough such that on expansion ofthe stent, the plaque or dissections of the body lumen are effectivelypressed back in position against the wall of the lumen giving adequatesupport to prevent tissue prolapse. At the same time, these elementsshould not be so close as to affect flexibility adversely, interferewith each other during crimping of the stent on the balloon of acatheter or exhibit inadequate access to the side branch in the vascularlumen. The design should be stiff enough to impart required radial andfatigue resistance strengths to the stent. The elements should undergoenough plastic deformation on expansion at specified pressure such thatthe elastic recoil is within acceptable limits. When the stent isexpanded radially, its diameter increases which causes change in itslength. The shape and arrangement of the elements should compensate thischange in stent length to maintain the original length as far aspossible keeping it within acceptable limits. This is achieved bycausing the specific strut elements to elongate in unison with radialexpansion. Though different sections of the stent may have differentmechanical strength across its axis, the stent should achieve itsspecified diameter uniformly across its entire length when the rateddeployment pressure is applied to the balloon catheter. The designshould offer adequate grip of the crimped stent on the balloon of thecatheter to resist dislodgement during delivery and minimum recoil afterexpansion of the stent at diseased site of the body lumen. The scaffoldstructure should allow re-intervention when the stent is implanted in alumen which has a side branch i.e. it should offer adequate side branchaccess. In such a case, the structural cells should create sufficientlylarge opening without breaking the struts when another catheter with orwithout a stent is inserted through the struts of implanted stent. Thestructure should have adequate strength and flexibility to withstand allforces of crimping on the balloon of the catheter, maneuvering throughthe vascular lumens, expansion/deployment at the diseased site andcyclic forces induced by the vascular lumen.

The stent has rows of cylindrical elements or struts which form rings.These cylindrical elements are interconnected by cross linking elementsor struts. The shape formed within the two consecutive cylindricalelements and two consecutive cross linking elements forms a ‘cell’ or‘cell structure’. The way such elements are interconnected with eachother can be manipulated to achieve different structural properties.Design flexibility is achieved by making these cells with varyinglengths and widths. Cells with larger length and width will give lowerstrength. On the other hand, cells with shorter length and width willgive higher strength. For same strut thickness, the struts with higherwidth will have higher strength and offer more resistance to compressiveforces of the arterial wall compared to struts of less width. The terms‘element’ and ‘strut’ are used interchangeably throughout thisspecification.

The scaffold structure of the stent of present invention generallyconsists of multiple rows, of sinusoidal wave type cylindrical elementswith regular or irregular shapes with plurality of peaks and valleysacross its axial length of the stent. The cells are formed by connectingupper and lower rows of cylindrical elements with straight or curvedlinks (“cross linking elements” or “cross linking struts”). These crosslinking elements connect upper and lower rows of cylindrical elementsanywhere along the length of the sides of the elements. Theseinterconnections form cylindrical scaffold structure of the stent.

The cross linking struts provide flexibility to the stent for easymaneuvering of the stent in curved and tortuous paths of the body lumen.The structural strength of the irregular curvilinear cylindrical elementcan be changed by changing the location where the cross linking strutsare attached along the length of the element. In the embodimentsdescribed in this invention, these linkages are located at peaks andvalleys or nearly at the center of respective sides of the elements. Thewidth and shape of individual strut and the cell are designed such as toprovide effective crimping, to impart sufficient radial strength inexpanded state and at the same time to keep recoil and change in lengthwithin acceptable limits. The scaffold structure after expansion givesacceptable side branch access. The irregular curvilinear line structurehas varying degrees of curvature in regions of the peaks and valleys.The curvature can be varied to impart different structural strength. Theshape should give uniform and low crimped profile as well as uniformradial expansion of individual elements around the circumference of thestent in a section and in individual layers along the axis of the stent.When rated deployment pressure is applied to the stent through theballoon of the catheter, the stent attains a uniform diameter across itsentire length in spite of having differential strength of elementsaxially.

The sinusoidal scaffold structures are designed with struts and straightor curved connecting linkages to give segments which are highlyflexible. On expansion of the stent during its deployment, thesesegments deform circumferentially from crimped diameter to an enlargedexpanded diameter. Different radial expansion characteristics can beobtained by changing size, shape and cross-section of the sinusoidalelement and cross linking struts. In addition, the radial strength ofthe stent can be increased by increasing the number of cells in a row.Similarly, the strength of the cells can be increased by increasing thenumber and width of the cross linking struts. The location where thecross linking struts connect the upper and lower rows of cells can alsobe manipulated to increase the strength and the overall flexibility ofthe stent.

The shape of the cells can be changed by changing, the curvature oftheir sides. In a limit, they can be given shape of a straight line.Such changes can make a difference in overall strength of the cell andhence the strength of the row and stent structure as a whole.

The geometry of the interconnected scaffold structure of the stent is sodesigned that the elastic recoil and change in length of the stent onexpansion are kept within acceptable limits.

The number of rows of elements to from cells is dictated by overalllength of the stent. The number of cells in a row along thecircumference of the stent, defined as crowns, is dictated by thediameter of the stent and width of the cell. The number of crowns can bechanged keeping balance with crimping profile.

The overall configuration of the stent decides the radial strength,flexibility and fatigue resistance of the stent. The dimensions of eachcell and their spacing are adjusted close enough to prevent protrusionof the plaque or any part of the body lumen where the stent isimplanted. At the same time, these dimensions are adjusted to achievetrouble free crimping of the stent over the balloon of the catheterwithout compromising the flexibility of the stent. The spacing is alsoadjusted to give desired side branch access. This configuration givesuniform coverage of lumen wall with the stent struts after the stent isfully expanded. The stent gets nicely and firmly apposed in the bodylumen. During deployment, the individual elements of sections may bedisturbed slightly relative to adjacent cylindrical elements withoutdeforming the overall scaffold structure. After the stent is expanded,portions of the elements may slightly tip outwardly and embed in thevessel wall a little to position the stent properly in the body lumen.This aids in apposing the stent firmly in place after expansion.

The configuration of the individual cells, the cross linking elementsand their interconnections are designed to distribute the stressesduring crimping and expansion uniformly across the entire stent.

The interconnection of cells with each other is achieved by the crosslinking elements as described above. These linkages are connected eitherat peak or the valley of the element forming the sinusoidal wave typeshape of a cell. These linkages can also be connected nearly at thecenter of the side of the element forming the sinusoidal wave type shapeof a cell. This gives a structure which is in the form of awell-supported structural beam in which the unsupported length isreduced at the connection point of cross linking element like a crosslinked truss girder beam. The cross linking elements can also beconnected off-center to the side elements of the cells. This will dividethe unsupported length of this element in 3 sections. The unsupportedlength of these elements depends on the positions of these cross linkingelements. The elements of the cells undergo full plastic deformationafter expansion to keep elastic recoil well within acceptable limits.

The configuration of the stent scaffold structures described above giveenough leeway to a stent designer to vary the shapes and otherdimensions of the elements of the stent to effectively reduce thethickness of the stent struts imparting necessary radial strength to thestent structure and get desired fatigue resistance. As described above,it is a well accepted fact that reduced thickness of the stent reducesinjury to the walls of the body lumen.

Flexibility of the stent is decided by the thickness and number of crosslinking elements across the circumference of the stent as well as theirlocations. If the number of these connectors is reduced, some of thesinusoidal sections become free to give more flexibility to the stent.However, this will reduce the mechanical strength of the stent. Thus itis extremely important to strike a balance between the flexibility andthe strength to optimize overall properties of the stent.

The designs of stents described in specific embodiments of thisinvention are based on the principles described above and are generallyfor coronary vasculature. However, the configurations described in thisinvention can be varied to get different shapes of the stent such thatit is possible to make stents for other applications like cerebralvasculature, renal vasculature, peripheral vasculature etc. by strikingbalance between strength and flexibility depending on the function. Inthis way, the stent structure configuration described in this inventiongives enough flexibility to a stent designer to tailor the stent for anyapplication.

Using the above general principles and specific scaffold structuredesign, it was possible to make bioabsorbabale stents with strutthickness of 130 μm or less, preferably between 100 μm and 110 μm. Thusthe invention further encompasses balloon expandable stent made frombioabsorbable polymer with thin struts (strut thickness 130 μm or less,preferably 100-110 μm) with high fatigue and radial strength.

Typical stent scaffold structures are described below and shown in FIG.4 to FIG. 7. These structures are described as typical examples and onewith skill in this art will appreciate that the low strut thickness of130 μm or less can be obtained by other designs with similar featuresusing the principles described above.

The scaffold structure consists of curvilinear sinusoidal shaped rows ofstruts. These rows are interconnected with cross linking struts to formthe overall stent structure. The shape of the rows and the way the rowsare interconnected can be altered to get desired mechanical strength andother essential properties of the stent like flexibility (pushabilityand trackability), lumen to stent surface area ratio, desired sidebranch access, desired crimping profile etc.

Scaffold structures designed based on above general principles aredepicted in FIGS. 4, 4A, 4B, 4C, 5, 6 and 7. These structures withadequate radial and fatigue strengths can be made with strut thicknesslower than 130 μm using polymer and the process of instant invention.

FIG. 4 shows a stent in flattened configuration in vertical positionwith a preferred scaffold structure which consists of rows of curvedstruts 100 with sinusoidal wave like shape with peaks P and valleys Vwhich form rings. The terms “peak” and “valley” are relative and dependson positioning of the scaffold structure. With reference to FIG. 4, peakis the portion which rises vertically up and valley is the portion whichshows a dip. The rows of struts are aligned in such a way that the peaksof one row/ring face the valleys of the subsequent row/ring and viceversa. The rows/rings of wavy sinusoidally shaped struts 100 areinterconnected by cross linking struts 101 to form the stent. The crosslinking struts 101 connect peak of the lower row with the valley of theupper row. The placement of cross linking struts 101 is after leavingtwo subsequent peaks and valleys and this forms the cell 103. Thesestruts 101 impart mechanical strength and connectivity to the structure.The length of the cross linking struts at the ends of the stentstructure (104) is kept little longer than other such elements tofacilitate fixing of the radio-opaque markers 102. The stent made ofthis design using the polymer and process of instant inventiondemonstrated adequate mechanical strength viz. radial strength andfatigue strength required for coronary stent with 125 μm strutthickness. This structure also demonstrated adequate trackability,pushability, sufficiently large side branch access and other suchessential properties.

The struts 101 can be placed on every alternate peak and valley as shownin FIG. 4A. This will increase the number of cross linking struts 101 instructure depicted in FIG. 4A compared to the structure depicted in FIG.4. Increased number of these cross linking struts 101 will impart highermechanical strength to the scaffold compared to structure depicted inFIG. 4. Thus, the structure of FIG. 4A is stronger than that of FIG. 4.Hence, same mechanical strength can be achieved in structure shown inFIG. 4A with thinner struts (<125 μm thick) compared to the structureshown in FIG. 4. In the limit, the cross linking elements can beprovided at every peak and valley as shown in FIG. 4B. The strength ofthe stent will be maximum in this case but this design will compromiseother properties like ease of crimping, flexibility, side branch accessetc. Thus, one has to strike balance between strength and otherproperties.

A variation of the scaffold design depicted in FIG. 4 is shown in FIG.4C. The design shown in 4C is same as that in FIG. 4 except the cells atboth the ends of the stent structure are made short by connection withcross linking elements 101′ at each peak and valley. Thus, the cells atthe ends become mechanically stronger than the other cells. When thisstructure is expanded, closed cells at the ends offer more resistance toexpansion compared to cells in the central portion of the stent. Hence,the stent will tend to expand from the central portion earlier than theend portions. This will result in the central portion contacting thearterial wall before the end portions. This eliminates classical“dog-boning effect” where the end portions expand before the centralportion causing edge injury to the artery during implantation. The stentmade of this design using the polymer and process of instant inventiondemonstrated adequate mechanical strength viz. radial strength andfatigue strength required for coronary stent with 105 μm strutthickness.

A similar scaffold structure with different geometry and shape is shownin FIG. 5. This structure also contains rows of sinusoidal struts 105with different curvature than the struts depicted in FIG. 4. Theinterconnecting struts 106 are not straight but are slanted. The crosslinking struts 108 at the ends are kept straight and radio-opaquemarkers 107 are located on these struts. The shape of the cell 109 islittle different than the structure shown in FIG. 4. This structure willhave somewhat different mechanical properties than the structuredepicted in FIG. 4.

A still different scaffold structure is depicted in FIG. 6. The rows 111are wavy but not sinusoidal in shape. They have a specific shape ofdifferent design. The cross linking elements 112 are straight. The radioopaque markers 113 are fixed on the cross linking elements at the endsof the stent.

A yet different scaffold structure is depicted in FIG. 7. The rows 114are again not exactly sinusoidal in shape but they have a designdifferent than design shown in FIG. 4 and FIG. 6. In this design, therows of struts are aligned in such a way that the peaks of one row/ringface the peaks of the subsequent row/ring and valleys of one row/ringface the valleys of the subsequent row/ring. The rows/rings of struts114 are interconnected by cross linking struts 115 to form the stent.The cross linking struts 115 connect peaks of consecutive rows. Theplacement of cross linking struts 115 is at every alternate peak andthis forms the cell 116. The cross linking elements 115 are straight butlonger than those in FIG. 6. The radio opaque markers 117 are fixed onthe cross linking elements at the ends of the stent.

Each of the stent designs described above follow general patternsdescribed earlier but has different properties and strengths. Using theprinciples described above, one skilled in this art can generate anumber of alternate designs with desired characteristics.

The surface of the laser cut stent is cleaned using iso propyl alcohol(IPA) to remove surface defects.

The stent should exhibit enough radio opacity for ease of implantationprocedure. A polymer stent does not have adequate radio opacity tobecome visible in X-ray imaging. The visibility in X-ray imaging isachieved by providing radio opaque markers on the stent. Theradio-opaque markers help to locate the position of stent during andafter the deployment with the help of X-ray imaging. During theoperation of laser cutting the stent pattern on the tube, holes ordepots are cut in the cross linking elements located on proximal anddistal ends of the stent structure where radio opaque markers are fixed.The radio opaque markers are deposited in these holes or depots by usingtweezer with or without a vacuum pump which can generate vacuum of 10 to15 inch of mercury. Radio-opaque markers are made from radio opaquemetals which should be biocompatible and should not interfere with thetreatment site. Such metals include platinum, gold, tantalum etc. In apreferred embodiment, six platinum markers are fixed on the stent, threeat each end of the stent equi-spaced circumferentially at 120° to eachother. The shape of these markers is tri-axial ellipse as shown in FIG.8. This shape is clearly visible in the X-ray imaging. This arrangementgives a clear idea of the stent position as well as patency of the stentat the ends in two standard orthogonal views without help of OCT orIVUS.

The process of placing the markers in the hole or the depot issimplified. The marker may be pressed against the hole or depot by aflat tool under optical microscope or under magnifying glass until themarker gets firmly fixed at the center of the hole or in the depot. Abiocompatible adhesive may be used for better securement of a marker inthe hole or the depot. The biocompatible adhesive is selected from butnot limited to the compounds like polyester, polyamides, PEG, proteins,cellulose, starch and their mixtures. Suitable solvent is used formaking adhesive glue. The solvent should be volatile enough to getevaporated to avoid presence of residual solvent on the stent. Thissolvent is selected from but not limited to the compounds likechloroform, ethanol, water, acetone or their mixtures.

In yet another embodiment, the annealing of the laser cut stent withradio opaque markers is done at temperature between 90° C. and 120° C.preferably between 100° C. and 110° C. for a period varying from 30 minto 16 hours, preferably from 2 hours to 8 hours and more preferably from3 hours to 4 hours. Vacuum of up to 650 to 700 mm Hg (absolute pressureof 60-110 mm Hg) is applied to remove the monomers. The stent is thencooled to ambient temperature in 20 sec to 10 minutes preferably between30 sec to 2 minutes. The stent at this stage achieves adequatemechanical strength viz. radial strength and fatigue strength. Thecrystallinity at this stage is between 40% and 50%.

The surface of the stent is then cleaned using solvent like iso propylalcohol (IPA), chloroform, perchloroethylene (as such or diluted with asuitable solvent) or mixture thereof. Cleaning operation removes surfacedefects and makes the surface smooth. The process involves dipping theannealed stent scaffolds mounted on mandrels in the solvent mixture toclean the scaffold. Cleaning is achieved by rotating the stent in thesolvent mixture for up to maximum 10 minutes at ambient temperature. Thecleaned stent is dried under vacuum to remove residual solvent on thesurface. The process is controlled such that desired strut thickness isachieved in this operation.

In a further embodiment, the stent is coated with formulation oftherapeutic agent by spray coating method. The therapeutic agent may beantiproliferative drug/s formulated with a carrier that may allowrelease of the therapeutic agent in controlled manner. The therapeuticagent and the carrier may be dissolved in a suitable solvent tofacilitate spray coating process. The solvent may then be removed byevaporation under vacuum. In one embodiment, a formulation consisting ofSirolimus drug and PDLLA polymer as carrier in 50:50 w/w proportion isdissolved in a suitable solvent and the solution is used for coating.The solvent is selected from compounds like methylene chloride,chloroform, acetone, methanol and mixtures thereof. The uniform andsmooth coating is achieved by spraying the solution on the stent from aspray coating machine. The parameters for spray coating process need tobe accurately controlled. These parameters include distance betweenstent and spray gun tip, collate rotation, solution flow rate and inertgas pressure used for spraying.

In a preferred embodiment, the parameters are as under.

-   -   The distance between spray gun tip and stent may be 3 cm to 10        cm, more specifically 4 cm to 6 cm.    -   The collate rotation speed may be between 10 and 20 rpm    -   Spray gun oscillation may be between 30 and 60 per minute and        more closely between 35 and 55 per minute.    -   The inert gas is nitrogen at 1.5 to 2.5 kg/cm² pressure.    -   Flow rate of the solution is kept between 0.10 and 0.40 ml/min,        more specifically between 0.15 and 0.30 ml/min.    -   After coating process, the stent is kept under vacuum to remove        the solvent.

The stent (with or without coating) is then crimped on the balloon ofthe catheter. Crimping operation is very critical. It should not affectthe stent surface and cause no mechanical damage to the coating or tothe stent. Crimping parameters include diameter of stent after crimping,pressure of crimping, dwell time and temperature. Crimping operation canalter properties of the polymer and hence the properties of the stent.The stent length and balloon size affect the crimping operation. Foreffective crimping, the crimping parameters change with stent length andballoon size. In a preferred embodiment, crimping is done in 6 to 8stages with dwell time between 200 to 310 seconds. The crimpingtemperature is kept between 25° C. and 40° C.

Finally, the stent prepared according to the invention is sterilizedusing e-beam sterilization process. This method is commonly used for thesterilization of medical devices because e-beam radiation can providemuch higher dosing rate compared to gamma rays or X-rays which reducesthe exposure time which in turn reduces potential degradation of thepolymer. Another advantage is that the sterilization process leaves noresidue. The sterilization is conducted at ambient or lower than ambienttemperature to avoid temperature related degradation of polymer.

It is known that e-beam radiation causes degradation of polymer and thushas a significant effect on the average molecular weight M_(w) andaverage molecular mass M_(n) of the bioabsorbable polymers and hence itsmechanical properties. We studied the effect of e-beam radiation over awide range of doses varying from as low as 5 kGy to as high as 50 kGy.The objective was to study the effect of e-beam dose on M_(w) and M_(n).The reduction in M_(w) varied from as low as 23% for e-beam dose of 5kGy to as high as 58% for dose of 50 kGy. Reduction in M_(n) varied fromas low as 32% for e-beam dose of 5 kGy to 67% for dose of 50 kGy.

It is hence very essential to reduce the e-beam dose to minimize thepolymer degradation. The effect of e-beam dose on the degradation ofpolymer can be reduced to some extent by adding a stabilizer in thepolymer matrix. This stabilizer should be biocompatible and should notcreate any adverse clinical effect.

Normal dose of e-beam for effective sterilization is more than 20 kGy.The instant invention reduces the degradation by reducing the dose ofe-beam considerably without compromising effective sterilization andwithout use of a stabilizer. The sterilization process of instantinvention is done in two parts as described hereafter.

The stent system consists of components like stent, and deliverycatheter. The e-beam dose affects the polymer stent and not the othercomponents. All components of the stent system other than the stent aresterilized separately using either Ethylene Oxide (ETO) or e-beam. Thepolymer stent is then mounted on the sterilized catheter in cleanenvironment by crimping process. The assembly along with the crimpedstent is then subjected to e-beam sterilization at a temperature between15° C. and 25° C.

Using this process, the effective sterilization was achieved with e-beamdose lower than 15 kGy; preferably between 5 kGy and 12 kGy, which ismuch lower than the normal effective dose of more than 20 kGy. Thechange in average molecular weight M_(w) of the polymer aftersterilization varied from 23% to 40% depending on the dose level. Thismethod resulted in acceptable bio burden of less than 3 cfu andSterility Assurance Level (SAL) of “six log reduction” (10⁻⁶). Thesterilization did not have any significant effect on the opticalrotation or crystallinity.

Following cases exemplify the process of sterilization and its effect onmolecular weights of the polymer. These cases are described for purposeof explanation and they do not limit the invention in any way.

Case-1:

The stent with M_(w) of 378,240 and M_(n) of 211740, crystallinity ofaround 48% and radial strength around 31 N was subjected to the abovesterilization process. The catheter and other components were sterilizedusing ETO. The unsterilized stent was crimped on the balloon of thesterilized catheter in clean environment and the assembly was packed inappropriate manner for e-beam sterilization with e-beam dose of 10 kGyat 15° C. The dose resulted in effective sterilization. M_(w) and M_(n)after sterilization were 275940 and 131310 respectively, crystallinity47% and radial strength of 27 N. The reductions in M_(w) and M_(n) were27% and 38% respectively. PDI changed from 1.79 to 2.1. The specificrotation was −158°.

Case-2:

The stent with M_(w) of 360980 and M_(n) of 200550, crystallinity of 51%and radial strength around 32 N was subjected to the above sterilizationprocess. The catheter and other components were sterilized using ETO.The unsterilized stent was crimped on the balloon of the sterilizedcatheter in clean environment and the assembly was packed in appropriatemanner for e-beam sterilization. Effective sterilization was achievedwith e-beam dose of 6 kGy at 15° C. M_(w) and M_(n) after sterilizationwere 268930 and 132760 respectively, crystallinity 49% and radialstrength of 28 N. The reductions in M_(w) and M_(n) were 25.5% and 33.8%respectively. PDI changed from 1.8 to 2.03. The specific rotation was−161°.

Case-3:

The stent with M_(w) of 345460 and M_(n) 189900 was subjected to e-beamsterilization process. The e-beam dose was 25 kGy at 15° C. M_(w) andM_(n) after sterilization were 206250 and 91152 respectively. Thereductions in M_(w) and M_(n) were 40.3% and 52% respectively. PDIchanged from 1.82 to 2.26. The specific rotation was −158°.

Case-4:

The stent with M_(w) of 352670 and M_(n) 192750 was subjected to e-beamsterilization process. The e-beam dose was 45 kGy at 15° C. M_(w) andM_(n) after sterilization were 165750 and 71320 respectively. Thereductions in M_(w) and M_(n) were 53% and 63% respectively. PDI changedfrom 1.83 to 2.32. The specific rotation was −154°.

EXAMPLES

The following examples are only for illustrating and help understandingthe invention. They do not limit the invention in any manner.

Example-1

The starting material for making the stent was extruded PLLA tube withM_(w) of 591280, M_(n) of 354890, PDI 1.67, specific rotation of −158°and glass transition temperature 60° C., and crystallinity 9%.

The tube was deformed at 74° C. by applying axial force till desiredstretch was achieved to get axial expansion ratio of 1.6. The conditionswere maintained for 15-20 sec and then the axial force was removed. Theradial expansion was then carried out by pressurizing the tube withnitrogen in 3 stages at 74° C. to achieve radial deformation ratio of 4.

Stage-1: 270 psi. Stage-2: 390 psi. Stage-3: 520 psi.

At each stage, the conditions were maintained for 15-20 sec.

While maintaining the pressure, the tube was heated to 110° C. overabout 1 min and this temperature was maintained for 1 min. The tube wasthen cooled to 20° C. over about 30 sec. The pressure was then releasedand the tube was removed from the mold.

The M_(w), M_(n) and PDI of the deformed tube at the end were 518350(reduction of 12.3%), 324760 (reduction of 8.5%) and 1.596 respectively.It is evident that the molecular weight distribution was narrower afterprocessing which is advantageous. (Single stage processing of the sametube at 150 psi pressure to achieve same radial deformation ratio, i.e.4, resulted into 16.2% reduction in M_(w), 18.0% reduction in M_(n), andPDI 1.7).

The deformed tube was cleaned with iso propyl alcohol and then cut onlaser machine using laser beam of 1400 nm wave length with pattern ofFIG. 4C to form the stent.

The laser cut stent was annealed at 105° C. for 3.5 hours under vacuumof 700 mm Hg. The stent was then cooled to ambient temperature in 1minute. The M_(w), M_(n) and PDI of the processed stent were 447620(reduction of 13.6%), 248160 (reduction of 23.6%) and 1.804respectively. The crystallinity at this stage was 48%.

The stent was then cleaned by rotation for 10 min in perchloro ethyleneat ambient temperature. At the end of this operation, the strutthickness of 105 μm was achieved.

3 platinum radio opaque markers of the shape depicted in FIG. 8 werefixed at each end of the stent without use of adhesive in the holesformed during laser cutting operation.

The stent was then crimped at 35° C. under clean environment onpre-sterilized PTCA catheter in 8 stages and total dwell time of 250-270sec.

The stent system was effectively sterilized using e-beam dose of 6 kGyat 18° C. M_(w), M_(n) and PDI of sterilized stent were 332130(reduction of 25.8%), 163290 (reduction of 34.2%) and 2.03 respectively.

The stent demonstrated radial strength of 20-25 N depending on stentdimensions and adequate fatigue strength.

Example-2

The starting material for making the stent was extruded PLLA tube withM_(w) of 605440, M_(n) of 366920, PDI 1.65, glass transition temperature62° C., crystallinity 11.5%, and optical rotation of −159 0.2°.

The tube was deformed at 75° C. by applying axial force till desiredstretch was achieved to get axial expansion ratio of 1.5. The conditionswere maintained for 15-20 sec and then the axial force was removed. Theradial expansion was then carried out by pressurizing the tube withnitrogen in 3 stages at 75° C. to achieve radial deformation ratio of3.9.

Stage-1: 280 psi. Stage-2: 400 psi. Stage-3: 510 psi.

At each stage, the conditions were maintained for 15-20 sec.

While maintaining the pressure, the tube was heated to 100° C. over 1min and this temperature was maintained for 1 min. The tube was cooledto 20° C. over 30 sec and pressure was released. The tube was thenremoved from the mold.

The M_(w), M_(n) and PDI of the deformed tube at the end were 526460(reduction of 13.05%), 326980 (reduction of 10.9%) and 1.61respectively. It is evident that the molecular weight distribution wasnarrower after processing which is advantageous.

The deformed tube was cleaned with iso propyl alcohol and then cut onlaser machine using laser beam of 1500 nm wave length with pattern ofFIG. 4C to form the stent.

The laser cut stent was annealed at temperature at 105° C. for 3.5 hoursunder vacuum of 700 mm Hg. The stent was then cooled to ambienttemperature in 1 minute. The M_(w), M_(n) and PDI of the processed stentwere 450120 (reduction of 14.5%), 248460 (reduction of 24.0%) and 1.812respectively. The crystallinity at this stage was 46%.

The stent was then cleaned by rotation for 10 min in perchloro ethyleneat ambient temperature. At the end of this operation, the strutthickness was 110 μm was achieved.

3 platinum radio opaque markers of the shape depicted in FIG. 8 werefixed at each end of the stent without use of adhesive in the holesformed during laser cutting operation.

The stent was then crimped under clean environment on pre-sterilizedPTCA catheter at 38° C. in 7 stages and total dwell time of 240-260 sec.

The stent system was effectively sterilized using e-beam dose of 10 kGyat 18° C. M_(w), M_(n) and PDI of sterilized stent were 321830(reduction of 28.5), 161240 (reduction of 35.1%) and 2.0 respectively.

The stent demonstrated radial strength of 20-25 N depending on stentdimensions and adequate fatigue strength.

1. A process for preparation of a biodegradable polymer stent made ofPLLA (poly-L-Lactide) with a strut thickness of less than 130 μm,comprising: (a) Deforming an extruded PLLA tube axially at a temperatureof 70° C. to 80° C. by applying an axial force until a desired stretchis achieved and then radially expanding the tube at a temperature of 70°C. to 80° C. by pressurizing the tube with inert gas in three stages,including pressurizing at a pressure of 250-280 psi in a first stage, apressure of 375-410 psi in a second stage, and a pressure of 500-530 psiin a third stage; (b) Heating the tube to a temperature between 100° C.and 110° C. after radial expansion under pressure for up to 2 min, andthen cooling the tube to 20° C. in 20-30 sec to get a finished deformedtube; (c) Cutting a scaffold structure pattern on the deformed tube bylaser machining to obtain a stent; (d) Annealing the laser machinedstent and depositing radio opaque markers on the stent; (e) Cleaning theannealed stent with radio opaque markers using a solvent to removeirregularities and to achieve a smooth surface; (f) optionally coatingthe cleaned stent with a formulation of an antiproliferative drug and acarrier polymer by spray coating; (g) Crimping the cleaned stent on aballoon of a pre-sterilized delivery catheter in a clean environment toform a crimped stent system; and (h) Sterilizing the crimped stentsystem by an e-beam method with an e-beam dose less than 20 kGy.
 2. Theprocess for preparation of biodegradable polymer stent made of PLLAaccording to claim 1, wherein the extruded PLLA tube is made of PLLAhaving a M_(w) in the range of 590000 to 620000, a M_(n) in the range of350000 to 370000, and crystallinity in the range of 7% to 12%.
 3. Theprocess according to claim 1, wherein the axial deformation of the tubeis carried out at temperature between 74° C. and 76° C. at axialexpansion ratio between 1.4 and 1.7 and maintaining the temperature andpressure conditions for 15-20 sec.
 4. The process according to claim 1,wherein the radial deformation of the tube is carried out at temperaturebetween 74° C. and 76° C. at radial deformation ratio between 3 and 5 bypressurizing the tube with nitrogen in three stages and maintaining thetemperature and pressure conditions for 15-20 sec at each pressurizationstage.
 5. The process according to claim 1, wherein the tube afterradial deformation, while maintaining the pressure, is heated between100° C. and 110° C. and maintained for 30 sec to 2 min and then cooledto 20° C. in 20-30 sec to get finished deformed tube.
 6. The processaccording to claim 1, wherein the laser cutting operation of thefinished deformed tube is done at wave length between 1300 and 1600 nm.7. The process according to claim 1, wherein the scaffold structure ofthe stent comprises a pattern with rows of struts with sinusoidal shapewherein the peaks of one row is aligned with valleys of the subsequentrow and wherein, the valley of one row is connected by straightcrosslinking struts with peak of subsequent row at every third suchposition in the central portion and at every peak and valley in the endportions.
 8. The process according to claim 1, wherein the annealing isdone at temperature between 100° C. and 110° C. for a period varyingfrom 3 hours to 4 hours under vacuum of up to 650 to 700 mm Hg followedby cooling the stent to ambient temperature in 25-30 sec.
 9. The processaccording to claim 1, wherein the cleaning of the annealed stent is doneby rotation in perchloro ethylene as such or diluted in a suitablesolvent or in a mixture of isopropyl alcohol and chloroform.
 10. Theprocess according to claim 1, wherein six radio opaque markers made fromplatinum with triaxial ellipse shape are fixed on the cross linkingstruts, wherein the said markers are equi spaced at 120° to each other.11. The process according to claim 1, wherein the stent is coated with aformulation comprising Sirolimus and a biodegradable polymer in a 50:50proportion w/w, at a Sirolimus dose of 1.25 μg per mm² of the stentsurface area.
 12. The process according to claim 11, wherein the coatingis done using spray coating.
 13. The process according to claim 1,wherein the cleaned stent is crimped on the pre-sterilized balloon ofthe delivery catheter under clean atmosphere at 25° C. to 40° C. in 6-8stages and dwell time of 200-310 sec.
 14. The process according to claim1, wherein the stent is sterilized using process steps as under: a)subjecting the components other than the stent to ETO sterilizationprocess; b) crimping the un-sterilized stent on a balloon of thepre-sterilized catheter in a clean environment; and c) subjecting thecrimped stent system to e-beam sterilization at dose between 6 and 12kGy, preferably between 6 and 10 kGy at temperature between 15° C. and25° C.
 15. The process according to claim 14, wherein the sterilizationprocess results in a bio burden of less than 3 cfu and a SterilityAssurance Level (SAL) of 10⁻⁶ with no effect on the optical rotation orcrystallinity.
 16. The process according to claim 1, wherein the stenthas a strut thickness of 100-110 μm and a radial strength of 20-25N. 17.A balloon expandable stent made from a bioabsorbable polymer by theprocess of claim 1, said stent having a strut thickness of 100-130 μmand a radial strength of 20-25N.
 18. The process according to claim 1,wherein step (d) comprises annealing the laser machined stent prior todepositing the radio opaque markers on the stent.
 19. The processaccording to claim 1, wherein step (d) comprises annealing the lasermachined stent after depositing the radio opaque markers on the stent.20. A process for preparation of a biodegradable polymer stent made ofPLLA (poly-L-Lactide), comprising: (a) deforming an extruded PLLA tubeaxially at a temperature of 70° C. to 80° C. by applying an axial forceuntil a desired stretch is achieved, and then radially expanding thetube at a temperature of 70° C. to 80° C. by pressurizing the tube withan inert gas; (b) cutting a scaffold structure pattern on the deformedtube by laser machining to obtain a stent having a strut thickness ofless than 130 μm; (c) annealing the laser machined stent; (d) optionallycoating the annealed stent with a formulation of an antiproliferativedrug and a carrier polymer; (e) crimping the annealed stent on a balloonof a pre-sterilized delivery catheter in a clean environment to form acrimped stent system; and (f) sterilizing the crimped stent system by ane-beam method with an e-beam dose less than 20 kGy.